Cone-beam tomography (CT) involves the imaging of the internal structure of an object by collecting several projection images (“projections”) in a single scan operation (“scan”), and is widely used in the medical field to view the internal structure of selected portions of the human body. Typically, several two-dimensional projections are made of the object, and a three-dimensional representation of the object is constructed from the projections using various tomographic reconstruction methods. From the three-dimensional image, conventional CT slices (e.g., cross-sections) through the object can be generated. The two-dimensional projections are typically created by transmitting radiation from a “point source” through the object, which will absorb some of the radiation based on its size and density, and collecting the non-absorbed radiation onto a two-dimensional imaging device, or imager, which comprises an array of pixel detectors (simply called “pixels”). Such a system is shown in FIG. 1. Typically, the point source and the center of the two-dimensional imaging device lie on a common line, which may be called the projection line. The source's radiation generally propagates toward the imaging device in a volume of space defined by a right-circular cone having its vertex at the point source and its base at the imaging device. For this reason, the radiation is often called cone-beam (CB) radiation. A full cone may be projected onto the imaging device, or a half cone may be projected, in which case the center of the imaging device is generally offset from the projection line (the configuration is generally called the half-fan configuration). Generally, when no object is present within the cone, the distribution of radiation is substantially uniform along any perimeter of any circle on the imaging device, with the circle being centered about the projection line and within the cone. However, the distribution of the radiation may be slightly non-uniform among the circle perimeters. In any event, any variation in the distribution can be measured in a calibration step and accounted for.
In an ideal imaging system, rays of radiation travel along respective straight-line transmission paths from the source, through the object, and then to respective pixel detectors without generating scattered rays. However, in real systems, when a quantum of radiation is absorbed by a portion of the object, one or more scattered rays are often generated that deviate from the transmission path of the incident radiation. These scattered rays are often received by surrounding pixel detectors that are not located on the transmission path that the initial quantum of radiation was transmitted on, thereby creating errors in the electrical signals of the surrounding pixel detectors. The collection of scattered rays from all the scattering events that occur during the collection of a projection is referred to herein as the scattered radiation.
The scattered radiation causes artifacts (e.g., phantom images) and loss of resolution and contrast in the CT image slices produced by the radiation imaging system. The size and magnitude of artifacts are generally larger in the half-fan configurations, where the center of the imaging device is offset from the projection line, than in the full-fan configurations. The scattered radiation can also cause numerical errors in the image reconstruction algorithms (generally referred to as “CT number problems” in the art). The foregoing lead to image degradation. Accordingly, there is a need in the cone-beam tomography field to reduce the impacts of scattered radiation.